The field of the invention is nuclear magnetic resonance (NMR) spectroscopy, and in particular, methods for localizing NMR chemical shift spectra of chemical moieties in the body produced during in vivo phosphorus-31, hydrogen-1 (.sup.1 H), and carbon-13 NMR spectroscopy.
Localized NMR spectroscopy of human organs in vivo has many clinical applications. For example, important metabolites contain phosphorus and the concentration or ratios of these metabolites provide valuable functional information. Abnormal ratios of phosphorus metabolites such as phosphocreatine (PCr), phosphodiesters (PDE), phosphomonoesters (PME) and adenosine triphosphate (ATP) have been observed in several human tumors and can be used not only to detect the tumors, but also, to measure the affect of chemotherapy. Also, the ratio of certain metabolite concentrations has been used to detect severe coronary artery disease without invasive diagnostic measures and before detectable pain is produced. As the instrumentation and methods improve, the measurement of various biochemical compounds and metabolites with spectroscopy will most certainly provide many additional clinical diagnostic benefits
The in vivo measurement of metabolite concentrations using NMR presents a number of challenges. Metabolites are at very low concentrations within the human body as compared with water, and as a result, metabolites will never be imaged with anatomical detail comparable to .sup.1 H images of tissue water. Signals from metabolites must be acquired over a substantial volume of tissue in order to produce an acceptable signal-to-noise ratio. Thus, while the signals may be "localized" to a particular organ or part of an organ, the localized volume from which the signal is acquired is substantially larger than the millimeter-sized voxels common to .sup.1 H imaging of water. Another difficulty in NMR spectroscopy is the very short spin-spin (T.sub.2) relaxation times of some of the chemical moieties. This means that the already small NMR signal produced by these moieties decay rapidly after the application of the RF excitation pulse. Consequently, it is important that the NMR signal be read out as soon as possible after termination of the RF excitation pulse. Therefore, the localization method used with spectroscopy must not lengthen the pulse sequence such that the NMR signal has decayed to an unacceptable low level by the time it is acquired.
One commonly used method for localizing the NMR signal in NMR spectroscopy studies employs a surface-coil that is positioned on the patient, immediately adjacent to the organ of interest. This method relies on the restricted range of sensitivity of the surface-coil to provide localization in the dimensions transverse to the central axis of the coil. Depth localization along the central axis of the coil is typically achieved by a one-dimensional phase encoding pulse which is produced during the pulse sequence. While the local coil is clearly more sensitive to NMR signals produced along its central axis, the surface-coil does pick up NMR signals from the surrounding region, particularly at greater depths. Thus, while the surface coil may be most sensitive to NMR signals produced by the organ of interest, the NMR signals from surrounding tissues may contribute enough to the total signal "seen" by the surface-coil to substantially distort the metabolite concentration measurements.
Such measurement errors can be partially corrected by employing a one-dimensional slice selective RF excitation pulse in each NMR pulse sequence. For example, if the surface-coil lies in the X-Z plane, a G.sub.z gradient may be applied concurrently with the RF excitation pulse to excite a slab of spins that include the organ of interest. Spins lying outside the excited slab along the Z axis will not contribute to the acquired NMR signal, and will not, therefore, distort the metabolite measurement. When combined with a G.sub.y phase encoding gradient applied during the pulse sequence, the region of interest is localized along 2 of 3 dimensions. While such two-dimensional localization is adequate for some measurements, it is not adequate for others.
More recently, three-dimensional localization has been achieved in NMR spectroscopy by using a so called two-dimensional selective excitation. In contrast to the well-known one-dimensional, slice selective, excitation which employs a constant magnetic field gradient during the application of the RF excitation pulse, two-dimensional selective excitation is achieved by applying two, orthogonal, time varying magnetic field gradients concurrently with the RF excitation pulse. In the above example, a G.sub.z gradient and G.sub.x gradient are applied during each excitation pulse. As described in U.S. Pat. No. 4,812,760 entitled "Multi-Dimensional Selective NMR Excitation With A Single RF Pulse", the time variations in the two orthogonal gradients and the amplitude envelope of the concurrent RF excitation pulse can be chosen to provide different shaped excitation patterns. In general, these two-dimensional selective excitation pulses are shaped to produce a cylindrical volume of excited spins located directly beneath the surface-coil and concentric with its central axis. Phase encoding is then used to locate spins along the length of the cylinder and to thereby provide the third dimension of localization.
A more recent variation of the two-dimensional selective excitation pulse is disclosed in an article by John Pauly et al. entitled "k Space Analysis Of Small-tip-angle Excitation", Journal of Magnetic Resonance 81, 43-56 (1989). This method excites a cylindrical region by producing two orthogonal gradients which vary sinusoidally and diminish to zero during the application of the RF excitation pulse. These gradient waveforms are illustrated in FIG. 1A and the amplitude envelope of the concurrent RF excitation pulse is illustrated in FIG. 1B. This method of two-dimensional selective excitation is referred to in the art as "spiral two-dimensional selective excitation" because the vector sum of the two applied orthogonal gradients map out a spiral pattern in k-space. The spiral pattern is illustrated in FIG. 3 and the resulting two-dimensional NMR sensitivity profile is illustrated in FIG. 1C. One of the added benefits of the spiral k-space trajectory is that the spins are not only transversely magnetized in the desired cylindrical-shaped region of interest, but they are also rephased. As a result, a subsequent rephasing gradient lobe is not required in the pulse sequence and the NMR signal can be acquired sooner with less T.sub.2 decay.
There are two problems which have heretofore limited the use of two-dimensional selective excitation in phosphorus-31 and carbon-13 spectroscopy and the like. The first limitation is the rate at which the amplitude of the magnetic field gradients can be changed. The gradient coils on whole body NMR scanners have substantial inductance and their slew rate is limited. As a result, to perform the spiral excitation illustrated in FIG. 1 within the slew rate limits of the NMR system, the RF excitation pulse may have to extend up to 20 milliseconds in duration resulting in substantial loss of signal from short T.sub.2 components.
Another limitation of prior methods using two-dimensional selective excitation pulses is the reduced bandwidth of the excitation. The bandwidth of the NMR excitation is inversely proportional to the duration of the selective excitation pulse. As indicated above, because of slew rate limitations in the gradient field generators, such pulses become very lengthy. The net result is a reduction in the breadth of the spectrum that is excited Referring to FIG. 2, for example, a phosphorus-31 spectrum showing a number of important metabolites requires a bandwidth of 30 ppm if the relative concentrations of all these metabolites is to be measured. If the two-dimensional excitation pulse requires eight times as long to complete, the useful spectrum that is excited will be significantly reduced and may not contain all, or even two, of the metabolite peaks that are to be compared.